Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc.
There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a bodily region of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis.
In traditional PET imaging, a patient is injected with a radioactive substance with a short decay time. As the substance undergoes positron emission decay, it emits positrons which, when they collide with electrons in the patient's tissue emit two simultaneous gamma rays. The gamma rays emerge from the patient's body at substantially opposite directions. These rays eventually reach a scintillation device positioned around the patient. There is often a ring of scintillation devices surrounding the patient. When the gamma rays interact with oppositely positioned scintillation devices, light is emitted and detected. The light is usually transmitted through a lightguide to a photodetector. The light detected by the photodetector is then interpreted by a processor to enable an image of a slice of the region of interest to be reconstructed.
In PET (as well as SPECT) it is important to match the scintillator emission wavelength to the photodetector's optimal wavelength quantum efficiency (QE). For example, a typical photomultiplier tube (PMT) used in PET applications has a peak wavelength sensitivity at 420 nm while a typical LSO scintillator used in PET emits at 420 nm. Therefore, PMTs and LSO are very well matched in terms of wavelength matching. LSO is a very good scintillator for a PMT but it is not well matched for use with other silicon-based photodetectors such as avalanche photodiodes (APDs) and silicon photomultipliers (SiPMs). These silicon photodetectors usually have a peak wavelength sensitivity at ≧500 nm. In some devices, such as SiPMs, their QE may increase 2-3 times from 420 nm to >500 nm. It is difficult to make a scintillator with good PET properties and to make it emit at an exact wavelength for photodetection. Such a scintillator may be made from crystal materials such as, but not limited to, LSO, YSO, LYSO, LuAP (i.e., LuAlO3:Ce), LuYAP, or LaBr3.
The phoswich approach has been used to improve the detection in PET applications by determining the depth-of-interaction (DOI) in the detector. PET scanners are typically made of long, thin detectors with high stopping power to meet high sensitivity requirements. In the absence of DOI information, however, the thickness of the scintillator reduces the spatial resolution due to parallax error. To compensate for reduced spatial resolution, detectors with DOI capability have been used. DOI capability can determine the location of the gamma interaction in the direction of the incident gamma (i.e., depth from the surface of the detector). One way to implement DOI capability is to use a multi-layer detector, in which the layers are made of material with different scintillation properties. Because the layers have different characteristics, when a gamma event is detected it is possible to identify which layer absorbed the gamma photon and so to determine more accurately the spatial interaction location in three dimensions. A “phoswich” thus is a detector with two or more layers of different scintillators. Layer identification is done by using differences in scintillation decay time and pulse shape discrimination techniques. The advantages of the concept have been demonstrated in the HRRT high resolution PET system using a LSO/LYSO combination giving a high spatial resolution uniformity of around 2.5 mm within a larger part of the imaged volume. Since LSO and LYSO have the same excitation and emission characteristics, layer identification is based on pulse shape discrimination as the LSO and LYSO materials have a difference in scintillation time of approximately 10-15 ns.
Two phoswich combinations that have been gaining popularity are LuAP/LSO and LuYAP/LSO. LaBr3/LSO has also been investigated. However, the functionality of these phoswich combinations is somewhat limited because both LuAP and LaBr3 emit in the excitation band of LSO, YSO or LYSO. For the phoswich to work well the scintillations of the two components must be independent of each other. However, since the emission of LuAP and LaBr3 scintillators are in the excitation band of LSO, both scintillators will be activated by a single 511 keV absorption and it is thus difficult to achieve a unique scintillation layer identification.
There thus remains a need in the art for a true phoswich (i.e., a phoswich using different scintillator types) that can provide an easily obtained unique detector identification for DOI calculations.